Growth factors (GFs) are peptides and proteins that stimulate the growth and/or differentiation of cells via the interaction of the GFs with specific cell surface receptors. Growth factors play an integral role in the repair and regeneration of tissues and exogenous application of GFs can be used to stimulate the repair of various tissues and organs including bone, cartilage, skin and mucosa and to enhance repair of tissues through the stimulation of angiogenesis at the repair site.
The transforming growth factor beta (TGFβ) superfamily of secreted growth and differentiation factors in mammals has over 30 members. These dimeric proteins are characterized by a conserved seven cystine knot-based structure. They regulate the proliferation, differentiation and migration of many cell types, and have important roles in morphogenesis, organogenesis, tissue maintenance and wound healing. The TGFβ superfamily of growth factors can be subdivided into several subfamilies including the transforming growth factor beta subfamily, the bone morphogenetic protein (BMP) and growth and differentiation factor (GDF) family (also called the BMP subfamily), and the inhibin and activin subfamily.
The BMP subfamily of the TGFβ superfamily comprises at least twenty proteins, including BMP-2, BMP-3 (also known as osteogenin), BMP-3b (also known as growth and differentiation factor 10, GDF-10), BMP-4, BMP-5, BMP-6, BMP-7 (also known as osteogenic protein-1, OP-1), BMP-8 (also known as osteogenic protein-2, OP-2), BMP-9, BMP-10, BMP-11 (also known as growth and differentiation factor 8, GDF-8, or myostatin), BMP-12 (also known as growth and differentiation factor 7, GDF-7), BMP-13 (also known as growth and differentiation factor 6, GDF-6), BMP-14 (also known as growth and differentiation factor 5, GDF-5), and BMP-15 (for a review, see e.g., Azari et al. Expert Opin Invest Drugs 2001; 10:1677-1686).
BMPs have been shown to stimulate matrix synthesis in chondroblasts; stimulate alkaline phosphatase activity and collagen synthesis in osteoblasts, induce the differentiation of early mesenchymal progenitors into osteogenic cells (osteoinduction), regulate chemotaxis of monocytes and mesenchymal cells, and regulate the differentiation of neural cells (for a review, see e.g., Azari et al. Expert Opin Invest Drugs 2001; 10:1677-1686 and Hoffman et al. Appl. Microbiol. Biotech 2001; 57:294-308).
One of the many functions of BMP proteins is to induce cartilage, bone, and connective tissue formation in vertebrates. The most osteoinductive members of the BMP subfamily are BMP-2, BMP-4, BMP-6, BMP-7, BMP-8 and BMP-9 (see, e.g., Hoffman et al., Appl. Microbiol Biotech 2001, 57-294-308; Yeh et al., J Cellular Biochem., 2005; 95-173-188; and Boden, Orthopaedic Nursing 2005, 24:49-52). This osteoinductive capacity of BMPs has long been considered very promising for a variety of therapeutic and clinical applications, including fracture repair; spine fusion; treatment of skeletal diseases, regeneration of skull, mandibular, and bone defects; and in oral and dental applications such as dentogenesis and cementogenesis during regeneration of periodontal wounds, extraction socket grafting, alveolar ridge augmentation, and sinus augmentation. Currently, recombinant human BMP-2 sold as INFUSE® by Medtronic FDA approved for use in spinal fusion surgery, for repair of fracture non-unions and for use in oral surgery, while and recombinant human BMP-7 sold as OP-1@ by Stryker is approved as an alternative to autograft in recalcitrant long bone nonunion and for revision posterolateral (intertransverse) lumbar spine fusions, where autograft and bone marrow harvest are not feasible or are not expected to promote fusion.
Other recombinant growth factors that have been used exogenously to enhance bone repair include various TGFβs (see Clokie & Bell, J. Craniofacial Surg. 2003, 14:268-77), members of the fibroblast growth factor superfamily (FGFs) (see Kawaguchi et al., (2007) J. Orthopaedic Res. 25(4): 480-487), members of the platelet derived growth factor superfamily (PDGFs) (see Hollinger et al., 2008 JBJS 90(s1):48-54), and vascular endothelial growth factor (VEGF) (Street et al., 2002 PNAS 99:9656-61).
For these growth factors to be effective they must be active and available at a sufficient concentration at the time when critical densities of the appropriate responsive cells are present in the repair site. The short half-life, thermal instability, sensitivity to proteases and/or solubility of the GFs requires their administration in combination with a carrier to achieve this requirement.
A number of carriers have been evaluated for the delivery of GFs. These include fibrous collagen sponges, gelatin hydrogels, fibrin gels, heparin, reverse phase polymers such as the poloxamers, carriers composed of poly-lactic acid (PLA), poly-glycolic acid (PGA) or their co-polymers (PLGA), heparin-conjugated PLGA carriers, and inorganic materials such as calcium phosphates. For example the bioimplant (GEM-21S®) which is used for periodontal regeneration uses beta tricalcium phosphate (β-TCP) as the carrier for rhPDGF-BB.
However, these carriers are of limited effectiveness, due to loss of growth factor activity when associated with the carrier, inefficient release of the GF at the implantation site, and/or poor protection from proteolysis and degradation. For example the bioimplant Infuse® uses a type I collagen sponge as the carrier for rhBMP-2. The rhBMP-2 is released in a burst from the carrier and the half life of the BMP within the wound site is 1-3 days (Winn et al., 1998, Adv. Drug Del. Rev. 31:303; Friess et. al., 1999, Intl. J. Pharm., 187:91). By the time the mesenchymal stem cells which regenerate bone have migrated into the wound site only fractions of a percent of the original amount of BMP loaded is present to stimulate these cells to make bone. The current solution to ensure an effective level of BMP remaining at these later times is to significantly increase the amount of BMP that is initially loaded. These increased doses increase the risk of complications including bone formation beyond the implant site, autoimmune responses and potentially cancer. Further this dramatically increases the cost of the implant.
Therefore, a need exists in the art for materials and methods which release growth factors with a profile which minimizes the amount of growth factor that needs to be loaded to achieve the required therapeutic effect.
One strategy is to encapsulate the GF in a biodegradable polymeric matrix that releases the GF with a sustained release profile over many days. For example BMPs have been combined with poly-lactic acid (PLA) or poly-lactic co-glycolic acid (PLGA) to produce sustained release profiles. However the incorporation of the BMP in the PLA or PLGA can denature the BMP reducing its activity and it is difficult to manipulate the release profile to optimize the effectiveness of the bioimplant. Further the degradation rate of these carriers is typically such that large amounts of GF remain locked away long after healing is complete. Consequently large amounts of GF need to be loaded into these polymers to ensure sufficient GF is present at the appropriate times.
Another strategy is to chemically immobilize the GF directly onto the surface of carrier. However this may result in partial or complete loss of activity of the GF, and may restrict the GF activity such that only those cells directly in contact with the carrier are able to interact with the GF and respond (see Steinmuller-Nethl, D. et al., Biomaterials, 2006, 27: 4547-56) which could be undesirable as the effect could be limited to the immediate interface with the carrier and not throughout the wound site.
The composition of the carrier can influence delivery of the GF. Calcium sulphate has been considered desirable as a bone substitute and GF carrier because it is osteoconductive, biodegradable, biocompatible and nontoxic (Chen et al., J. Craniofacial Surg., 2010, 21:188-197). However, calcium sulphate is also known to have a rapid degradation rate when added to bone in situ and little osteoinductive capability, which has limited its usefulness in bone implants.
One strategy to manage calcium sulphate degradation in situ has been to control degradation rate by altering crystal structures and adding polymers (e.g., chitosan) to the calcium sulphate implant mixture (Chen et al., supra). Polymer-coating calcium sulphate pellets that have been impregnated with BMP can decrease the speed of resorption of calcium sulphate and increase compressive strength and osteoinduction of the mixture (Chen et al., supra).
Composites containing hydroxyapatite (HAp), a major mineral component of bone, and calcium sulphate hemihydrate (CSH, plaster of Paris) have been used in orthopedic grafts (e.g., Damien, C et al., J. Biomed. Mat. Res., 1990, 24: 639-654; Damien, C et al., Spine, 2002, 16S: S50-S58; Parsons, J., et al., Annals N.Y. Acad. Sci.). When CSH is mixed with sterile saline or water it immediately begins to gel. While in the gel state HAp, growth factors and/or various matrix components can be mixed together with the CSH to form the graft composite, which can be inserted or injected into a bone defect where it sets in situ. In such methods, CSH initially acts as a binder. However, subsequent resorption of calcium sulphate leaves behind a porous matrix with space for bone in-growth, which can be stimulated by the growth factors in the hardened composite. Similarly, compositions for delivering osteogenic proteins including CSH, a porous particulate polymer mixture and an ostogenic protein are known (U.S. Pat. No. 5,385,887 and U.S. Patent Application Publication No. 2008/0233165, each of which is incorporated herein by reference as if set forth in its entirety). In each of these methods calcium sulphate degradation is required for growth factor release. Therefore, bone regeneration is dependent on the rate of calcium sulphate degradation.
Bone grafts containing particulate bone and a biocompatible solid component comprising CSH and a calcium phosphate product are known, but do not involve using the CSH or calcium phosphate as a growth factor carrier (U.S. Patent Application Publication No. 2011/0208305, incorporated herein by reference as if set forth in its entirety).
In nature during wound healing multiple GFs are present within the wound site and surrounding tissue at varying concentrations at different times. For example, immediately following bone fracture, platelets at the injury site will initially release large amounts of PDGF, with a sharp decline in protein levels within the fracture site over the following days (see Tyndall et al., Clinical Orthopedics and Related Research, 2003, 408: 319-330). Conversely BMP-2 is expressed at all stages of the fracture healing process (see Rasubala et al. British Journal of Oral and Maxillofacial Surgery, 2003, 41: 173-178), although the amount of BMP-2 varies over time (see Meyer et al. J Bone Jt. Surg 2003, 85-A: 1243-1254). The concentration of these growth factors is estimated to be orders of magnitude lower than those used during therapeutic application of exogenous GF due to matching of the concentration to the cellular requirements and synergistic effects of the multiple growth factors. Producing a system that allows the delivery of growth factors with multiphasic release profiles and the release of multiple growth factors with different release profiles would permit the use of bioimplants with GF release profiles that more closely mimic GF release during the natural healing process than current bioimplants that release a single growth factor in a burst or with sustained release.
This background information is provided for the purpose of making known information believed by the applicant to be of possible relevance to the present invention. No admission is necessarily intended, nor should be construed, that any of the preceding information constitutes prior art against the present invention.